Method for producing strain induced austenite

ABSTRACT

This disclosure relates to shape memory alloys which have been subjected to a thermal and mechanical treatment to increase the austenite start temperature A s →A s ′ such that the shape memory alloy is martensitic at body temperature and when subsequently subjected to a controlled deformation, the shape memory alloy preferentially reverts to the parent phase. One application for this disclosure is a stent for use in a lumen in a human or animal body having a generally tubular body formed from a shape memory alloy which has been subjected to a thermal and mechanical treatment so it deforms as martensite until a critical expansion diameter is reached at which point the tubular body rapidly reverts to the parent phase with much higher mechanical properties. The shape memory alloy comprises Ni—Ti and a ternary element ranging from about 3 at. % to about 20 at. %. The ternary element is effectively insoluble in a Ni—Ti matrix. In a preferred embodiment, the element is selected from the group consisting of niobium, tantalum and zirconium.

CROSS REFERENCES TO RELATED APPLICATIONS

This application is a continuation-in-part of U.S. patent applicationSer. No. 11/909,681, filed Sep. 25, 2007, now U.S. Pat. No. 7,988,722B2, issued Aug. 2, 2011, which claims the benefit of U.S. ProvisionalApplication No. 60/846,951, filed Sep. 25, 2006, and PCT/US2006/010555,filed Mar. 24, 2006, which claims the benefit of U.S. ProvisionalApplication No. 60/665,526, filed Mar. 25, 2005, and each of which isincorporated herein by reference in its entirety.

TECHNICAL FIELD

The present disclosure relates to the processing of nickel-titanium(Ni—Ti) ternary alloys and, in particular, to the conditioning of thesealloys in a manner that enhances the use of such alloys in medicalapplications.

BACKGROUND

Metallic engineering materials such as stainless steels, cobalt chromiumalloys (Co—Cr alloys) and nickel-titanium alloys (Ni—Ti alloys) are usedin variety of medical device applications. One example of a Ni—Ti alloyis NITINOL. The use of these alloys combine mechanical, fatigue,corrosion resistance and biocompatibility properties to create devicesuseful in a number of medical procedures.

Since the advent of minimally invasive procedures, engineering designershave been trying to work with specific geometries whereby metalliccomponents can be inserted through very small openings in the humanbody, routed to the desired location and then deployed to a useful sizeto fulfill the needs of the end application. One such well-known deviceis a coronary stent, a tubular structure used to hold open blocked orcollapsed arteries. The usual method of getting stents into the space isto collapse the metal structure onto a delivery catheter having asufficiently small overall diameter so it can be routed percutaneouslyto the coronary artery and expanded to a much larger diameter than theoriginal insertion diameter.

Conventional alloys used in various medical instruments have relied onstainless steel, complex cobalt chrome alloys (such as Elgiloy™ or L605)all of which can have their mechanical properties (i.e. yield strength,ultimate tensile strength break strength, etc) modified through workhardening and annealing. These metals, even with very high yieldstrength, cannot sustain strains much greater than 0.2 percent (%)without suffering a permanent set. Once a bend or kink has beensustained in a medical instrument or device fabricated from one of theabove alloys it is virtually impossible to straighten and remove fromthe body. For many permanent implants (such as stents), the device maynot need to be removed and the permanent deformation may actually beuseful to keep the structure in place. However, in the foregoing alloys,Hook's taw dictates that the force to deploy the implant will increaseat a linear rate until the material yield point is reached and then theforce will continue to increase until the material break point isreached. Additionally, these materials have significant spring backafter receiving a significant deformation.

Recently medical device engineers have begun designing metalliccomponents with shape memory alloys. In general, shape memory alloyssuch as NITINOL having the proper transformation temperature andprocessing could potentially offer two modes of shape recovery formetallic components inserted into the human body: (1) superelasticityand (2) shape memory recovery.

In the case of superelastic NITINOL the complete “elastic” recovery ofstrains up to 10% due to stress induced martensite (SIM) can beachieved. When superelastic NITINOL components are subjected to astress, the strain is accommodated by austenite to martensitecrystalline transformation, rather than by the mechanisms that prevailin other alloys such as slip, grain boundary sliding and dislocationmotion.

Under typical process condition the stress required to form martensitewill be >60,000 pounds per square inch (psi) (414 mega pascals (MPa) andthe reverse transformation stress will be >30,000 psi (207 MPa). It canbe observed that the reversion stress is lower than the stress at whichmartensite forms. These stresses are referred to as the upper and lowerplateau stresses and their magnitude is dependent on the alloycomposition, cold working and thermal treatment that the NITINOL, hasreceived. As the temperature of the specimen is raised, the stressmagnitude required to produce SIM is increased; however when thespecimen reaches a critical temperature above (the Austenite finishtemperature) A_(f), designated as M_(d), stress induced martensitecannot be formed, no matter how high the stress. In practicalapplications, this behavior gives rise to a limitation on using thesuper-elastic property since it limits the temperature range over whichsuper-elasticity is observed; typically in the binary Ni—Ti alloys, thisis a temperature range of about 60° Celsius (C) (108° Fahrenheit) (F),although a 40° C. (72° F.) range is more typical. The desirabletemperature range for medical and orthodontic applications is in theregion of body temperature, +10° C. to +40″C, can be achieved in thesealloys.

Others have applied superelastic NITINOL to medical devices using a 50.8atomic percent (at. %) nickel/balance titanium formulation which hasbeen cold worked followed by a low temperature anneal to give acombination of shape memory and/or superelastic characteristics. Forstarting materials having an ingot A_(f) of 0° C., this processing givesa component with an elastic range of approximately 2% to 8% over atemperature range of +15° C. to +40° C. However, the shortcomings fordeployment of a stent application include: (a) the unnecessary bulk ofthe stent delivery system (since the delivery system must resist thehigh outward radial force of the compressed stent during shipping,storage and deployment); (b) the high outward radial force of thecompressed stent pressing on the inside surface of the delivery sheathcan add unwanted friction during deployment of the stent from thesheath; (c) at deployment the rapid stent expansion to its memorizedshape can traumatize the vessel wall; and (c) the stent can cause achronic outward force once deployed that can cause further trauma.

Jervis, U.S. Pat. No. 5,067,957, discloses that a medical devicecomponent made from superelastic NITINOL can be externally constrainedoutside the body via the stress induced martensite mechanism, thenplaced in the body and de-constrained for deployment.

Duerig, et al., U.S. Pat. No. 6,312,455, discloses a superelasticNITINOL stent for use in a lumen in a human or animal body, having agenerally tubular body formed from a shape memory alloy which has beentreated so that it exhibits enhanced elastic properties with a point ofinflection in the stress-strain curve on loading. This enables the bodyto be deformed inwardly to a transversely compressed configuration forinsertion into the lumen and then revert towards its initialconfiguration, into contact with and to support the lumen. The shapememory alloy comprises nickel, titanium and from about 3 at. % to about20 at. %, of the alloy composition, of a ternary element selected fromthe group consisting of niobium, hafnium, tantalum, tungsten and gold.The ratio of the stress on loading to the stress on unloading at therespective inflection points on the loading and unloading curves is atleast about 2.5:1, and the difference between the stresses on loadingand unloading at the inflection points is at least about 250 MPa.

Besselink, et al., U.S. Pat. No. 6,428,634, discloses a method ofprocessing a highly elastic stent made from a Ni—Ti—Nb based alloy whichcontains from about 4 to about 14 at. % Nb and in which the atomicpercent ratio Ni to Ti is from about 3.8 to 1.2, comprising working thealloy sufficiently to impart a textured structure to the alloy, at atemperature below the recrystallization temperature of the alloy.Preferably, the alloy is worked at least 10%, by a technique such asrolling or drawing, or another technique which produces a similarcrystal structure. The alloy has increased stiffness compared with Ni—Tibinary alloys with superelastic properties.

For the case of shape memory recovery mentioned earlier, thethermoelastic shape memory alloys can change from martensite toaustenite and back again on heating and cooling over a very smalltemperature range, typically from 18° C. to 55° C. On cooling from theaustenitic phase, often called the parent phase, martensite starts toform at a temperature designated as M_(s) (martensite start) and uponreaching the tower temperature, M_(f) (martensite finish), the alloy iscompletely martensitic. Upon heating from below the M_(f) temperaturethe martensite starts to revert to the austenitic structure at A_(s),and when the temperature designated as A_(f) is reached, the alloy iscompletely austenitic. These two crystalline phases have very differentmechanical properties: the Young's Modulus of austenite is 12×10⁶ psi(82,728 MPa), while that for martensite is about 4×10⁶ psi (27576 MPa);and the yield strength, which depends on the amount of cold work thealloy is given, ranges from 28 to 100 thousand pound per square inch(ksi) (193 to 689 MPa) for austenite and from 10 to 20 ksi (68 to 138MPa) for martensite.

Additionally, a NITINOL structure processed to exhibit shape memory anddeformed in the martensitic state can recover up to 8% strain on heatingto austenite. This would be an extremely handy way to deploy devices orrecover accidental bending and kinking of devices in the human body ifit were not for the heating and cooling extremes that must be achieved.

Simpson, et al., U.S. Pat. No. 4,770,725, discloses a Ni—Ti—Nb shapememory alloy and article, wherein niobium varies from about 2.5 to 30at. %. Also disclosed is an article made from thesenickel/titanium/niobium alloys.

Simpson, et al., U.S. Pat. No. 4,631,094, discloses a method ofprocessing a nickel/titanium-based shape memory alloy. The methodcomprises over deforming the alloy so as to cause at least some amountof non-recoverable strain, temporarily expanding the transformationhysteresis by raising the austenite transformation temperature, removingthe applied stress and then storing the alloy at a temperature less thanthe new austenite transition temperature. Simpson also discloses anarticle produced from this method.

Wu, et al., U.S. Pat. No. 6,053,992, discloses a mechanism that uses theshape recovery of a shape memory alloy for sealing openings orhigh-pressure passages. A component made of a shape memory alloy can beprocessed in its martensitic state to have a reduced dimension smallerthan that of the opening or the passage to be sealed. Upon heating,shape recovery takes place that is associated with the reversecrystalline phase transformation of martensite. The shape recovery ofthe previously processed shape memory alloy component yields a diametergreater than that of the opening or passage to be sealed. The shaperecovery provides the dimensional interference and force required forsealing.

Wu, et al., builds on the work of Simpson, et al., U.S. Pat. No.4,770,725, to use both the defined chemistry and the specified processmethod to effect a specific heat sealing application which employs aheat activated recovery transformation.

The wide thermal hysteresis available from thermal and mechanicaltreatment of alloys disclosed in the literature is attractive forarticles which make use of a thermally induced configuration change,since it enables an article to be stored in the deformed configurationin the martensite phase, at the same temperature at which it will thenbe in use, in the austenite phase. This thermal and mechanical treatmentis used in a variety of industrial heat—to recover couplings andconnectors (L. Mcd. Schetky, The Applications of Constrained RecoveryShape Memory Devices for Connectors, Sealing and Clamping, ProceedingsSuper-elastic Technologies, Pacific Grove, Calif. (1994)).

It has been reported that a reverse transformation start temperatureA_(s)′ has been raised to +70° C. after specimens were deformed to 16%strain at different temperatures, where the initial states of thespecimens were pure austenite phase and/or martensite phase dependingupon the pre-straining temperature regime. It was found that atransformation hysteresis width of 200° C. could be attained and thereverse transformation temperatures were measured by forcing ashape-memory recovery via heating, and that up to 50% of the pre-straincould be recovered. The work was done by Xiang-Ming He, et, al, Study ofthe Ni_(41.3)Ti_(38.7)Nb₂₀ Wide Transformation Hysteresis Shape MemoryAlloy, Metallurgical and Materials Transactions, Vol. 35A, September2004. The work cited is an optimization of various pre-strainingconditions to maximize the strain recovery possible for heat recoveryapplication.

While the wide hysteresis confers certain advantages when the thermallyinduced changes in configuration are to be exploited, a wide hysteresisin stress-strain behavior is generally inconsistent with the propertiesof an alloy that are desirable in stent or medical device applications.

Various methods have been described to deliver and implant stents. Onemethod frequently described for delivering a stent to a desiredintraluminal location includes mounting the expandable stent on anexpandable member, such as a balloon, provided on the distal end of anintravascular catheter, advancing the catheter to the desired locationwithin the patient's body lumen, inflating the balloon on the catheterto expand the stent into a permanent expanded condition and thendeflating the balloon and removing the catheter. One of the difficultiesencountered using other stents involved maintaining the radial rigidityneeded to hold open a body lumen while at the same time maintaining thelongitudinal flexibility of the stent to facilitate its delivery.

What has been needed and heretofore unavailable is a stent which has ahigh degree of flexibility so that it can be advanced through tortuouspassageways, can be readily expanded, and yet have the mechanicalstrength to hold open the body lumen into which it expanded.

Thus, it is desirable to develop an alloy that is very ductile anduniquely suited for deployment of medical devices, such as stents, intothe human body. In particular, for stents, it is desirable that thecompressed stent maintain its shape until expanded.

SUMMARY

In one embodiment; the present disclosure provides a Ni—Ti ternary alloythat is particularly useful for medical instruments and devices, as wellas components thereof.

In another embodiment, the present disclosure provides an alloy havingvariable hardness/stiffness properties and which is useful for medicalinstruments and devices, as well as components thereof.

In yet another embodiment, the present disclosure provides a materialfor making medical instruments and devices as well as components thereofthat are formable without crack initiation sites when expanded ordeformed.

Another embodiment is directed to a method of producing a Ni—Ti ternaryshape memory alloy, specifically strained induced austenite (SIA), withimproved characteristics which are desirable for use in variousapplications such as medical devices.

The disclosure provides for a stress induced martensite that is lockedplace by the presence of a third element such that A_(s)′>37° C. Thematerial can be placed inside a mammalian body that can have itsA_(s)′→A_(s) by controlled deformation, thus creating a material havingmechanical properties superior to the mechanical properties uponinsertion. When the material is used in a medical device structure, thedevice can be placed inside a mammalian body and restored to a stiffermechanical state after controlled deformation or expansion.

The material is particularly advantageous in that a medical devicecomprising the material can be placed inside a mammalian body that canhave its A_(s)′ initially reduced to a temperature <37° C. by controlleddeformation, and further restoration of the original A_(s) andaustenitic transformation of mechanical properties by heat driven shapememory recovery process. Once placed inside a mammalian body, thematerial can be restored to a stiffer mechanical state after controlleddeformation whereby A_(s)′<37° C., and the austenitic transformation iscompleted by a combination of controlled deformation and assisted bytemperature driven shape memory recovery process.

The disclosure is also directed to a method for the thermal mechanicaltreatment of a Ni—Ti alloy performed at temperatures <M_(d) (thetemperature at which martensite can no longer be stress induced) wherethe hysteresis is widened such that after the treatment is completedA_(s)<A_(s)′, or when used for medical applications, A_(s)<37°C.<A_(s)′.

The present disclosure provides a method of producing a Ni—Ti ternaryalloy that exhibits properties desirable for medical instruments anddevices. In particular, the treated alloy exhibits a high degree ofductility and low mechanical strength properties during insertion intothe body and subsequently can be made stiff with significantly highermechanical properties after being subjected to a controlled deformationof a critical magnitude.

The present disclosure is directed to a method for the thermalmechanical treatment of a Ni—Ti ternary alloy performed at temperatures<M_(d) wherein the hysteresis is widened such that after the treatmentis completed A_(s)<A_(s)′, comprising:

-   -   a) annealing or partially annealing a Ni—Ti ternary alloy        through a high temperature solution treatment followed by        cooling;    -   b) mechanically straining the alloy be under a load while        simultaneously cooling the material to a temperature around or        less than M_(s), thus shifting A_(s) up much higher such that        A_(s)<A_(s)′ and retaining a sufficient amount of strain in the        strained element whereby if controlled deformation is applied        the alloy is transformed to the austenitic phase, shifting        A_(s)′ to A_(s).

In preferred embodiments, the alloy is mechanically strained betweenabout 10 to about 25% under a constant load while simultaneously coolingthe material to a temperature around or less than M_(s), thus shiftingA_(s) up much higher such that A_(s)<A_(s)′ and retaining about 2% toabout 15% strain in the strained element whereby if controlleddeformation is applied the alloy is transformed to the austenitic phase,shifting A_(s)′ to A_(s).

Mechanical straining of the alloy during cooling forms martensitevariants having volume fractions comprised mainly of (a) stress inducedmartensite and or (h) twinned and deformed martensite. The volumefractions formed are dependent upon the cooling rate and the appliedpre-strain load stress. For instance if the applied pre-strain is at atemperature >A_(f) and <M_(d), the microstructure may reveal formationof stress induced martensite and if the pre-strain is applied at atemperature <A_(f) the micro structure may reveal formation of twinnedand deformed martensite. Optimization of the martensite volume fractionsare dependent upon the process pre-straining conditions and temperature.Reorientation of the martensite variants are prevented by the presenceof a third insoluble element in the metal matrix giving rise to A_(f)′.

Controlled deformation is displacement and the associated stress/strainfield (bending, compression, tension, shear) required to initiate acomplete or partially complete transformation wherein A_(s)′ is shiftedback to A_(s). During controlled deformation the prepared material isdeformed easily at low stress levels, and after a sufficient range ofdisplacements have been completed, the material undergoes a permanentshift in mechanical properties such that further displacement now occursat higher stress levels.

By pre-straining the Ni—Ti ternary alloy using the method describedabove, it has been found possible to produce a shift of a normallyaustenitic alloy at body temperature, into an alloy that has martensiticproperties at body temperature.

Surprisingly, the controlled deformation causes the pre-strainedmaterial to become much stiffer where there appears to be almost aspontaneous conversion from martensite to austenite. The controlleddeformation appears to “unlock” the martensitic alloy structure andcause the mechanical properties to revert seemingly spontaneously backto the austenitic alloy. Adding controlled deformation in theappropriate amount will cause the A_(s)′ to be restored to the originalmaterial A_(s). In certain embodiments of the disclosure, the controlleddeformation may occur through bending, compression, tension and/or shearstresses.

The resulting material is martensitic at all temperatures <A_(s)′, veryductile and uniquely suited for deployment of medical devices such asstents into the human body. After pre-straining the material, thecompressed stent maintains its shape until balloon expanded.

In one aspect of the disclosure, the shape memory alloys processed bythe method comprise a ternary element E from 3 at. % to about 20 at. %of the alloy composition E can comprises a ternary element that iseffectively insoluble in the Ni—Ti matrix such as Nb, Ta or Zr, and thelike.

The pre-straining process, in combination with the preferred Ni—Ti alloycomposition employed renders the components flexible that, in turn, makemedical components, in particular stents, easy to insert and deliverinto the desired location. The resulting alloy is a material that canrevert from martensite to austenite by deformation process as describedabove in (b). This characteristic is particularly advantageous whentrying to create a structure in the human body because delivery can bethrough a small opening and then after deployment the structure takes onthe stiffer rigid property.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing and other Objects and features of the disclosure willbecome apparent from the following description of preferred embodimentsof the disclosure with reference to the accompanying drawings, in which:

FIG. 1 is a graph showing radial expansion force versus stent diameterfor a 316L or Co—Cr annealed stent material;

FIG. 2 is a graph of temperature vs. strain showing the shift ofA_(s)→A_(s)′ after an alloy material [1] has been processed by the firstpart of method while [2] is an example of an alloy according to thepresent disclosure, processed by the method;

FIGS. 3 a and 3 b shows two graphs comparing pre-strained Ni—Ti-E alloyheating (I) and deformation (II) recovery modes;

FIG. 4 is a schematic representation of an apparatus used to perform athree point bend test to show the onset of “unlocking” mechanism;

FIG. 5 is a graph of a three point bend test performed with theapparatus of FIG. 4. These plots show first and second bending cycles ofstrain induced austenite and the unloading and reloading which occur at0.060 thousandths (″) (0.1524 centimeter (cm) displacement giving someindication of the spring back characteristic;

FIG. 6 is a schematic representation of some components used in theAcculine bend moment tester used to provide cyclic bend testing with aconstant moment arm;

FIG. 7 is a graph of the test results of the tests run with theapparatus of FIG. 6 showing strain induced austenite undergoingstiffness transformation during its first and second bending cycle;

FIG. 8 is an apparatus to measure of the generic stent cell performancecharacteristic;

FIG. 9 is a graph of data of the test using the apparatus of FIG. 8;

FIGS. 10A-10C show an apparatus used fix a test that shows the materialtransformation occurring by cyclic bending of the Strain InducedAustenite (SIA) material;

FIG. 11 is graph of the data using the apparatus in FIGS. 10A-10C;

FIG. 12 is a graph showing stent diameter vs. radial expansion of astent prepared from the strain induced austenite; and

FIG. 13 shows the strain behavior for a stent fabricated from straininduced austenite in relation to FIG. 11.

DETAILED DESCRIPTION OF THE ILLUSTRATIVE EMBODIMENTS

Most stents on the market today are made from materials such as 316L orCo Cr (L605 and MP35N) materials. The stents are compressed, permanentlydeformed and crimped onto a delivery catheter equipped with anunderlying balloon. Once positioned in the lumen, a balloon is inflatedto outwardly expand the stent material beyond its yield point, until thestent makes contact with the vessel wall.

FIG. 1 shows how the radial expansion force varies as a function ofstent diameter. These stent materials have four limitations: (1) theexpansion force increases with expansion diameter (FIG. 1 [1]); (2)material spring back requires (FIG. 1 [2]) over expansion to achieve thefinal diameter and this results in vessel trauma; (3) the high force toexpand the stent requires an inflation balloon to be of a heavy wallthickness increasing the overall profile and stiffness of the deliverysystem; and (4) under X-Ray fluoroscopy some of the above mentionedmaterials have poor visibility.

Additionally, some stents on the market today are made from superelasticNITINOL. In these devices, the “active” A_(f) (as defined by ASTM F2082)of the final stent structure is between 0° C. and 37° C., moreparticularly between 10° C. and 25° C. These stents are radiallycompressed at (a) room temperature against a radial outward forcegenerated by the upper plateau stress (about 70 ksi or 483 MPa) ofsuperelastic NITINOL or (b) at a temperature below M_(s) against themechanical stress (about 12 ksi or 83 MPa) of martensite. In eithercase, at room temperature the outward radial force of the compressedstent diameter can be constrained with a delivery sheath fitting overthe compressed stent. These stent materials have the followinglimitations: (1) the constrained outward force of the stent requires anouter sheath which adds unwanted stiffness and bulk to the deliverysystem; (2) the continually outward force of the compressed stent exertsa frictional force on the sheath which increases the sheath retractionforce; (3) once the sheath is pulled back the stent rapidly expands tothe final diameter which may cause vessel trauma; (4) the stent diametercannot be over expanded to accommodate any gap between the vessel lumenand the selected stent; and (5) under X-Ray fluoroscopy, binary NITINOLhas poor visibility.

In one aspect of the disclosure, the shape memory alloys processed bythe method comprise a ternary element E from 3 at. % to about 20 at. %of the alloy composition, wherein E can comprise a ternary element thatis effectively insoluble in the Ni—Ti matrix. Examples of suitableternary elements include, but are not limited to, niobium (Nb), tantalum(Ta), Zirconium (Zr) and the like. It is also thought that shape memoryalloys comprising combinations of element E can be effective.

TABLE I Composition of Materials Suitable for “Strain Induced Austenite”Alloy Titanium Nickel E Composition (Atomic %) (Atomic %) (Atomic %)Ni—Ti-E Y % Z % 0.1 < X % < 20% 1. X is the percentage addition of athird insoluble element (E) such as Niobium (Nb), Tantalum (Ta) andother insoluble ternary elements into the Nickel Titanium matrix. Theelement is supplied in sufficient amounts to optimize the “locking andunlocking” mechanism of “strain induced austenite.” 2. Y and Z are usedto adjust the ratio of Nickel and Titanium necessary to achieve thetarget ingot A_(s) temperature considered optimal by the endapplication. The ratio of Atomic Percent Ni to Atomic Percent Ti is fromabout 1.6 to .60.

In particular, the insoluble ternary element (Nb, Ta, or Zr) appears toplay a role to “lock” the A_(s)′ structure during the pre-strainingcycle in such a way that restoration of the austenite state is preventeduntil either (a) the materials is heated above A_(s)′ and shape recoverytakes place or (b) the controlled deformation sets up a strain fieldsufficient to trigger material “unlocking” and restoration of originalmaterial.

In contrast with other materials, the alloy of present disclosure isparticularly well suited for structures that are packed in a smallconfiguration, easily expanded with minimal force, and later afterdeployment becomes significantly stiffer, a property that is desired forvarious medical components, particularly stents.

Example 1 Preparation of the Pre-Strained Alloy

The method was applied to Ni₄₄T₄₇Nb₉ alloy to provide the responseillustrated in FIG. 2 which showed an increase of A_(s)→A_(s)′ and thewidening of the hysteresis from [1] to [2]. The alloy was partially orfully annealed between 450° C. to 900° C., particularly 650° C. to 900°C. for <60 minutes wherein after cooling the austenite start temperatureA_(s) was about −65° C. The material was strained under load while thematerial was cooled, preferably to a temperature between the martensiticstart temperature of M_(s) about −85° C. and the A_(s) about −65° C. Thematerial was then strained 8% to 25%. Upon completion of thepre-straining cycle A_(s)′ was elevated above 37° C. and particularlyabove 60° C.

Another aspect of the disclosure is the effective insolubility ofniobium in the Ni—Ti matrix. It appears that during the treatmentmethod, there is a partitioning of the total strain of the material andeach component (i.e. nickel-titanium and beta niobium) seems to work asan individual component at similar levels of flow stress. Depending uponthe pre-straining temperature regime, martensitic Ni—Ti deformsreversibly by either twin boundary motion or stress induced martensitemotion, while the beta-niobium deforms irreversibly via slip to “lock”the martensitic deformations in place. In either case the lockedstructure mechanism corresponds to the new A_(s)′.

After treatment, the Ni—Ti—Nb microstructure is comprised of: (a) stressinduced twined martensite consisting of type I (111) M twins andoccasionally (001) M twins and anti-phase domain boundaries (Icomet-92);and (b) the soft beta-niobium particles are deformed irreversibly viaslip.

The plentiful block shaped niobium phases are adjacent to the differentmartensite variant boundary below M_(s). After pre-straining, theseblock phases serve as obstacle sources to inhibit the reorientation ofthe martensite, so martensite reorientation has difficulty proceeding atlower stress levels. When stress is added by controlled deformation(e.g., bending), the martensite reorientation is free to proceed oncethe blocking constraint of the niobium phases is removed by theadditional stress.

Another aspect of the method is the creation of expandable structuresthat can be collapsed and ready for small opening insertion, that do notrequired external constraints. In this way, the pre-strained structureis comprised of (a) stress induced martensite or (b) twinned anddeformed martensite which is “locked” in place by the present of thethird and insoluble element (in this case the soft niobium particles).This type of structural element may not need an external cover to holdback the restoring force typically found in superelastic NITINOLexpandable structures. One advantage to the method is that no externalcover is required for insertion, leading to a reduced design profilethat is advantageous for a number of medical device applications.

A tubular stent element can be manufactured by one of the followingseries (e.g., 1, 2 or 3) of process steps. Note that the thermalmechanical treatment can occur either before or after the machining stepin items 1 and 2 below.

1) Starting Tube—Laser or Chemical Machining—Pre-Straining to increaseA_(s)′—Deployment and Expansion

2) Starting Tube—Pre-Straining to Increase A_(s)′—Laser or ChemicalMachining—Deployment and Expansion

3) Starting Wire—Shape Setting—Pre-straining to IncreaseA_(s)′—Deployment and Expansion

Other appropriate manufacturing methods incorporating the treatmentmethod would be known to those skilled in the art for use inapplications with wire, tube, strip or appropriate forms.

As discussed above, the present disclosure offers an alternate method todeploy structures and devices in the human body that are made from othermaterials (i.e. 316L and Co—Cr alloy systems) and super elastic NITINOL.

When the pre-strained Ni₄₄T₄₇Nb₉ material is heated such that T>A_(s)′,the Ni—Ti matrix reverts to its parent phase (Table 1—Process I) with acorresponding increase of mechanical properties (FIG. 3—Process I)sufficient to overcome the “locking” force of the deformed beta niobiumparticles and thereby “unlocking” the structure and recovering both theoriginal pre-strained shape and austenite start temperature (A_(s)). Thedrawback to the resulting material is that heating is not a preferredmethod to deploy a medical device into a mammalian body.

TABLE I Illustrates the Process Differences Between Process I (Heating)and Process II (Controlled Deformation) which is Strain InducedAustenite for Ni₄₄ Ti₄₇ Nb₉. Process Method Initial PreparationPre-Straining Application Recovery Phase I Cold Working Pre-strain whileConstrained Heating T > A_(s)′ plus Annealing @ cooling article to aboutRecovery To achieve: A_(s)′ → A_(s) 450 to 850° C. M_(s) or < M_(s)Application (e.g. pipe and shape recovery 1 to 30 minutes temperature tocoupling) achieve: A_(s) → A_(s)′ II Cold Working Pre-strain whileBiomedical Device Controlled deformation plus Annealing cooling articleto about Insertion (e.g. stent) of sufficient scale to 450 to 850° C.M_(s) or < M_(s) achieve: A_(s)′ → A_(s) 1 to 30 minutes temperature toachieve: A_(s) → A_(s)′

in the present method, the pre-strained material is “unlocked” by acontrolled deformation (Table 1 Process II) and upon reaching a criticalstrain level (ε_(c)) (FIG. 3—Process II), the soft martensitic structurebegins to revert to the parent phase having improved mechanicalproperties useful in many medical devices including stents.

Specifically, a controlled deformation of sufficient strain (ε_(c)) cancause the parent phase to preferentially nucleate in a jump-like manner.At locations where flow stress levels are sufficiently large, softniobium particles no longer constrain martensite variants which are now“unlocked” an undergo a parent phase transformation as a series ofstress level reductions (FIG. 3—Process II).

Example 2 Three Point Bend Test

Referring to FIGS. 4 and 5, a three point bend test was performed todemonstrate the “unlocking” clicking phenomena, the spring backcharacteristic in the “unlocking region” at 0.060″ (0.15 cm)displacement, and to show the stiffness transformation between the firstand second bend cycle.

A 0.020″ (0.05 cm) 316LVM (low carbon vacuum melted) wire in the fullyannealed condition (typical of condition and chemistry for stentmaterial) was prepared.

For the Strain Induced Austenite (SIA), 0.020″ (0.05 cm) Ni—Ti—Nb waspre-strained under approximately 302 ksi (2082 MPa) during a coolingcycle from 20° C. to −140° C. The wire had about 10% retained strainafter returning to room temperature.

A test apparatus 10 for the above method is shown in FIG. 4 below. Wirespecimen 12 was placed across a 0.375″ (0.952 cm) unsupported span 14.An Instron 5544 equipped with a 20 pound (lb.) (9.072 kilogram (kg))load cell and Blue Hill software was used for data capture. Anvil 16 wassecured in the Instron cross head and (a) advanced down at 0.04″/minute(0.10 cm/min) for a distance of 0.060″ (0.15 cm) (b) moved upward untilthe load reached zero, and (c) advanced down until total displacementequaled 0.100″ (0.254 cm).

FIG. 5 shows the load displacement behavior for 0.020″ (0.05 cm) SIAwire 12, subsequent bending cycle after a strain induced austenitictransformation and a 316LVM fully annealed wire of equivalent crosssection. The “unlocking mechanism” of the SIA material 12 was apparentin the displacement range from about 0.03″ (0.076 cm) to 0.08″ (0.2032cm). The onset of material “unlocking” occurred at 0.03″ (076 cm) andthe load capacity was suddenly shifted from 0.28 lbs (127 grams (g)) to0.16 lbs (72 g). There were a number of other small drops in loadcapacity observed up to 0.080″ (0.203 cm) of displacement. The neteffect of the unlocking mechanism is that bend force for the SIAmaterial did not grow very large. In comparison, the performance of 316Lfully annealed material (with well documented extremely low stiffnessand ductile material properties) was perhaps stiffer by a factor of twothroughout the range of displacement.

At 0.060″ (0.15 cm) wire 12 was unloaded to zero and then the load wasreapplied. The spring back for the SIA material was about 0.015″ (0.038cm) and the 316LVM fully annealed material was about 0.009″ (0.0229 cm).The test was stopped at 0.93″ (2.36 cm) of displacement. The peak forcefor SIA was 0.33 lbs (0.15 kg) while the peak force for 316LVM was 0.62lbs (0.28 kg). The data demonstrates that: (a) the initial bendingstiffness of SIA is about one-half (½) that of annealed 316 LVM, (b) the“unlocking” phenomena is prevalent in the range between 0.03″ (0.0762cm) and 0.08″ (0.203 cm) of displacement and is responsible for keepingthe stiffness extremely low during initial bending. In addition, it iswell documented that fully annealed 316LVM has yield strength ofapproximately 45,000 psi (310 MPa), whereas the first bend cycle of SIAmaterial has a yield point considerably less than the 316LVM.

After the initial bending cycle the deformed SIA wire 12 wasstraightened and retested under the same 3 point bend test conditions.These results are shown in FIG. 5 and identified as SIA—Subsequent BendCycle. At about 0.06″ of deflection the subsequent bending cycle afterstrain induced austenitic transformation showed a wire 12 that was morethan twice as stiff as the 0.020″ 316LVM fully annealed wire and morethan 4 times stiffer than the 0.020″ SIA wire during the first bendingcycle. This is evidence of the bending induced transformation frommartensite to austenite.

Example 3 Observation of the Reverse Phase Transformation by SimpleBending

Further evidence of strain induced austenite was easily observed usingwires having a starting diameter of 0.020″ which, after thepre-straining sequence described above, resulted in stable martensitematerial having a diameter of 0.0192″. As described above, there are twopaths by which the original 0.020″ wire diameter can be recovered: (1)by simple bending and (2) by heating. The heat to recovery method forthese alloys has been successfully employed for industrial applicationsbut is not considered acceptable for medical applications because of thehigh temperature (e.g., greater than 60° C.) required to effect thetransformation. However, if the pre-strained wire is bent between onesfingers at room temperature, the transformation taking place can beobserved by a soft clicking action that ends in a stiff bent wire. Ifthe bent wire is now straightened by bending in the opposite directionmore soft clicking can be observed. The end result is a phasetransformation of wire from a soft martensitic material into a stifferaustenitic material. Proof that a reverse transformation has taken placecan be found by measuring wire diameter with calipers. After bending thepre-strained wire as described above, wire recovered in diameter from0.0192″ after pre-straining, to its original 0.020″ diameter. Thisgrowth in wire diameter is proof of the recovery transformation viabending.

Example 4 Acculine Bend Moment Testing

Referring to FIGS. 6 and 7, this test compared the bending momenttransformation behavior of 0.020″ pre-strained Ni—Ti—Nb as describedabove and 0.020″ superelastic straight binary NITINOL. The bendingstiffness of strain induced austenite and superelastic NITINOL straightwire samples were tested under identical conditions using an AcculineAE#-BM Bend Tester 110. Each 0.020″ wire specimen 112 was mountedbetween custom mounting blocks 114 with a 1 millimeter (mm) bend radius.Drive pins 116 and 118 rotated at 9.0 degrees per second (°/sec) tomaintain a constant 0.3 cm moment arm while wire 112 was deflected fromthe vertical position 0°) counter clockwise to +45°, returned to 0°,counter clockwise to −45° and back to 0° to complete one bend cycle.Data was collected by a 10 in-oz rotary torque measurement sensors 120and 122 and captured for graphical Microsoft Excel presentation. FIG. 6shows the experimental set up. FIG. 7 is a graph of the data from thefirst, second and third bend cycle. The first bend cycle shows a peakmoment for the Ni—Ti—Nb to be about 41% less then superelastic NITINOL,the second bend cycle shows the peak maximum bend moment to be about 13%of the peak moment of superelastic NITINOL, and the third cycle showsthe Ni—Ti—Nb is approaching the performance of the superelastic NITINOLwire. This is further evidence that pre-strained Ni—Ti—Nb had undergonea transformation from a locked NITINOL structure (A_(f)′) to an unlockedstructure restoring the original A_(f).

Example 5 Expansion of Two Parallel Wires

Referring to FIGS. 8 and 9, the behavior of the material duringexpansion was tested, specifically, 0.020″ (0.05 cm) 316LVM (low carbonvacuum melted) wire in the fully annealed condition (typical ofcondition and chemistry for stent material). SIA 0.020″ (0.05 cm)Ni—Ti-Nib pre-strained under approximately 302 ksi (2082 MPa) during acooling cycle from 20″C to −140″C. The wire had about 10% retainedstrain after returning to room temperature.

Referring to FIG. 8, the testing apparatus 210 comprised two smallDelran blocks 212 and 214 which were devised to hold two 0.020″ (0.05cm) wires 216 and 218 in a parallel starting configuration (wire spacing0.100″ (0.254 cm)). The spacing 220 between blocks 212 and 214 used toclamp the parallel wires 216 and 218 was about 1.414″ (3.59 cm). AnInstron 5544 equipped with a 20 lb (9.072 kg) load cell and Bluesoftware was used for data capture. Specially designed micro-liftinghooks 222 and 224 (0.080″ or 0.203 cm width) were designed to expand theparallel wires by hooking to the mid point of the 1.414″ (3.59 cm) spanand then traveling a total distance of 1″ (2.54 cm) at a rate of 0.25inch per minute C/min) (0.635 centimeter per minute (cm/min) whilerecording the load. Once the crosshead reached a travel distance of 1″(2.54 cm), the expanded structure was unloaded until the crossheadreached a travel distance of 0.96″ (2.44 cm).

As shown in FIG. 9, the load deflection behavior of 0.020″ (0.05 cm)Ni—Ti—Ni wire has compared with the 31.6LVM fully annealed wire ofequivalent cross section. The “unlocking mechanism” of the StrainInduced Austenite (SIA) material was plainly visible (fine jaggedpattern) in the extension range from about 1″ (0.254 cm) to 0.7″ (1.8cm). In this range there was negligible force increase. The resistanceto bending of the 0.020″ (0.05 cm) SIA was less than half that of the0.020″ (0.05 cm) 316LVM fully annealed material. It is quite well knownthat 316LVM in the fully annealed condition has a yield strength ofapproximately 45,000 psi (310 MPa) and this would lead us to state thatthe yield point of the SIA material on the first cycle was considerablyless than 316LVM, which is already an extremely soft ductile material.This curve also shows that the work required to achieve a givendisplacement was considerably reduced to achieve a given displacementwhen compared with 316LVM fully annealed material. After expansion to 1′(2.54 cm), the curves were unloaded to 0.96″ (2.43 cm) and the load wasrecorded. During unloading the slope of the spring back was calculatedand it was found that the 0.020″ (0.05 cm) SIA material was 9.058 poundsper inch (lb/in) (1.411 newtons per millimeter (N/mm) while the 316LVMfully annealed material was 16.6 lbs/in (2.907 N/mm).

Example 6 Material Transformation Through Bending of the SIA Material

Referring to FIGS. 10A-10C and 11, a test was constructed to show thematerial transformation occurring by bending of the SIA material.

SIA −0.020″ (0.05 cm) Ni—Ti—Nb pre-strained under approximately 302 ksi(2082 MPa) during a cooling cycle from 20° C. to −140° C. The wire hadabout 10% retained strain after returning to room temperature.

A straight length of 0.020″ (0.05 cm) SIA material specimen 310 wasclamped in mounting blocks 320 at one end and at a distance of 0.20″(0.5 cm) from the fixed end, a close fitting guide 330 machined from aplate (0.080″ or 0.2 cm width) was attached. The close fitting guide 330was attached to the crosshead 340 of an Instron 5544 equipped with a 20lb (9.072 kg) load cell and Blue Hill software for data capture. Usingthe crosshead extension, the test began at the neutral position “A”(zero deflection with zero load). The cross head cycled upward +0.1″(−0.25 cm) to position “B” and downwards −0.2″ (−0.51 cm) to Position“C” and then upward +0.2″ (+0.51 cm) to Position “B” and so forth untilthe test was completed. The cross head speed was 0.2″/min or about 0.5cm/min.

As shown in FIG. 11, start at (A) and progress along with low force withevidence of “unlocking” to (Point B). The material was bent in thereverse direction from (Pt B) to (Pt A) with no force (more unlocking)until it got back slightly past the origin at (Pt A). Continued bendingtowards (Pt C) caused rapidly increasing stiffness with the force at (PtC) more than twice the force at (Pt B). During the second round trip toPt B1 and Pt C1, the results confirmed the material transformation wascompleted, demonstrating that mechanical bending converted A_(s)′ toA_(s) for this material system (i.e., Process II of Table I). Thedetails are discussed below.

The results in FIG. 11 show the load deflection behavior of 0.020″ (0.19cm) SIA wire for negative and positive bending conditions.

Displacement A→B

The test began at Position A (zero deflection and load). During positivebending to Position B, there was evidence of the “unlocking mechanism”at about 0.02″ (0.5 cm) when the load was reduced to near zero level(shown by a large instantaneous spike). The load recovered and overallstiffness remained relatively low as the displacement increased to about0.07″ (0.17 cm) at which point the stiffness increased quickly.

Displacement B→C

During unloading, the load reduced quickly to a displacement of 0.075″(0.19 cm) as would be expected. Beyond this range, there was evidence ofadditional material “unlocking” as the specimen bent easily withoutresistance and with near zero stiffness in the displacement range of0.07″ (0.18 cm)→0.01″ (0.0254 cm). At the displacement of −0.01″ (0.0254cm), the “unlocking mechanism” appeared complete, and the specimen hasnow transformed into a much stiffer material as evidence by theincreasing load at displacement Position C.

Displacement C→B1

The material has been completely transformed and the load at B1 hasincreased by more than twice when compared with Position B.Additionally, in the displacement range from 0.0″ (0 cm)→0.05″ (0.127cm), the load during the second cycle was four (4) to six (6) timesgreater than on the first cycle. This material test shows that astiffness transformation has occurred and the “unlocking phenomena”observed in the first displacement cycle has disappeared.

Displacement B1→C1→A

Subsequent cycling shows the material is no longer exhibiting the“unlocking” phenomena and bent with considerably more stiffness thanduring the first cycle.

This type of curve certainly demonstrates that mechanical bendingconverted A_(s)′ to A_(s) for this material system (i.e., Process II ofTable I). The near zero stiffness during bending, followed by atransformation to a significantly stiffer wire is unique, and offers thepotential to engineer many useful devices.

A further advantage of the present disclosure is to provide a thermalmechanical treatment regime such that A_(s)′ is sufficiently >37° C.,such that little or no heat induced shape memory recovery occurs duringtemperature exposures caused by placement of the stent into the humanbody, curing of drug coatings, sterilization or shipping of finishedmedical devices.

In another aspect of the disclosure, it is possible to deploy a highlyelastic (super-elastic) stent without having to constrain andde-constrain stress induced martensitic stent structure in contrast to,for example, Jervis, U.S. Pat. No. 5,067,957, discussed above. In thisaspect of the disclosure, a Ni—Ti-E material (as described above) havingappropriate chemistry to yield a fully annealed temperature A_(s) about−15° C. and ideally processed to have super-elastic properties at bodytemperature (particularly having an “active” A_(f) between 10° C. and20° C.) can be “locked” by the thermal mechanical treatment means suchthat A_(s)′>37° C. When placed in the human body without constraint andsubsequently balloon expanded, the structure can “unlock” at which pointA_(f)′ again approaches A_(f). Applying the straining techniques above,such a material's A_(s) could be shifted higher and the resultingstructure would be martensite at body temperature. At this point,insertion into the body followed by the controlled deformation would besufficient to restore the original A_(s) and therefore the superelasticproperties at room temperature. This alternate deployment methodeliminates the force required to hold the stent in the collapsed state,reduces high friction loads between the stent and delivery sheath andminimizes the stent delivery profile.

By reverting the material back to austenite (as A_(s)′→A_(s)), thesuperelastic Ni—Ti-E properties are reached at a certain point and thenit can be further expanded easily when needed—a property particularlydesirable for stent material. Furthermore, the material would possessthe advantages of radiopacity and higher strength when compared tobinary NITINOL.

A stent fabricated from “strain induced austenite” as described in thisdisclosure can have a much different stent radial deployment force thanthose of other alloys or superelastic NITINOL. Referring to FIG. 12,initially the stent radial expansion force increases in a linearreversible fashion with stent diameter and corresponds to the elasticdeformation limit of martensite along the path from the origin→P. Theradial force required to achieve point P is much less than what isrequired to achieve the yield point of other materials (316L and Co—Cr).This is a particular advantage of the current disclosure. The“unlocking” process of martensite begins at point P and ends at point B.As the stent diameter increases, the required radial outward force isconstant or decreasing during the “unlocking” process P→B where thematerial completes a reverse transformation to austenite.

Advantageously, when the diameter of the “unlocked” stent is sizedappropriately to the vessel lumen (this would correspond to point B inFIG. 12), then no overexpansion of the lumen is required, and thusvessel trauma is eliminated or minimized. Furthermore if the vessel walltries to contract or collapse, the stent offers a reserve of radialresistive force, as shown FIG. 13, IV. The reserve of radial resistanceis coming from the material transformation of martensite to austenite bythe controlled deformation of the stent expansion cycle. The change inmaterial stiffness is shown by examining FIG. 5 SIA—subsequent bendingcycle.

In the present disclosure, the force required to reach the fullydeployed stent diameter (point B FIG. 12) can be much less than theexpansion force required by a stent made from other materials. Thereduced expansion force can lead to an optimization of the stentdelivery profile.

Biomedical Device Applications

In general, systems using the present material can provide highermechanical properties than other binary alloys (for example, NITINOL,resulting in smaller device cross sections and minimal design profile.Such devices can reduce trauma since they do not have to be overdeformedduring deployment, as in the case of materials such as stainless steel(316L), Co—Cr alloys (L605, MP35N), and titanium-based materials.Another advantage of the present materials is less inflation pressure ofthe balloon. The addition of Nb or Ta into the Ni—Ti-E alloy can improvethe radio-opaque properties of the material, allowing doctors to findthe location of smaller cross sections under X-Ray fluoroscopy. Thealloy exhibits nonmagnetic, low torque properties, and offers a crispimage under MRI imaging which is a medically desirable property.

For percutaneous, intraluminal and laproscopic medical deviceapplications, the present disclosure offers multiple advantagesincluding: very low deployment forces, delivery systems with moreflexible and smaller cross sections, and inflation balloons with thinnercross sections and lower operating pressures for safer and higherreliability. Designs using the present material do not have to be heldin the compressed position awaiting deployment, such as binary NITINOLat high stress levels (>60 ksi or 413 MPa) during shipping,sterilization and storage.

Stents and Stent Grafts

As discussed above, stents are fabricated from laser cut tubes, braided,coiled or formed wires fabricated into tubular structures and used torepair the patency of narrowed, previously weakened or ballooned andotherwise impaired lumen or other body channels. They are deployed bythe use of catheters in percutaneous, intraluminal or laproscopicprocedures. Examples are: blood vessels, bile duct, esophagus, urethra,trachea and the like. Specifically: carotid and coronary vessel,intraluminal lining of aortic abdominal aneurysms, iliac or femoralaneurysms, recanalization of injured vessels caused by blunt orpenetrating trauma, dilation and recanalization of stenotic arterialsegments, tampanade and obliteration of esophageal varices,recanalization of esophageal stenoses secondary to carcinoma or benignstrictures, ureteral strictures and tracheal strictures. In all theseapplications, the present shape memory alloy would be advantageous inits ease of deployment.

The present disclosure improves on the current state of the art inseveral ways and the specific advantages depend upon the base materialsystem in the comparison.

For example, the present disclosure improves over a NITINOL stent byhaving little or no outward radial force when placed in the deliverysystem tube. A binary NITINOL stent exerts a chronic outward force onthe inside wall of the delivery system, and during storage the insidewatt of the delivery system sheath may become imprinted by the stentframe. During deployment of the NITINOL stent, the frictional forces maybe quite high, whereas devices formed from the present alloys deploymore easily and provide a more flexible and reduced delivery sheathcross-section. Furthermore, after the expansion of a stent made from thepresent alloys is expanded, and the transformation described herein iscomplete, the new material can provide stiffer characteristics than, forexample, binary alloys such as NITINOL. A further advantage is thatniobium is very radiopaque under x-ray fluoroscopy whereas NITINOL isnot.

When compared with a 316L stent, the present disclosure reduces thedelivery system profile, the balloon expansion pressure is reduced andthe total amount of work required to deploy the stent system is alsoreduced. The expansion force of a 316L stent increases linearly as thestent diameter is expanded while the present disclosure can achievesignificant stent expansion at near zero force as the diameter expands.The lower expansion force characteristic leads to a reduced crosssection of the balloon inflation catheter, thereby leading to a lowerprofile and improved flexibility for the delivery system. A furtheradvantage is that niobium is very radiopaque under X-ray fluoroscopy,whereas 316L is considered to have poor visibility.

FIG. 13 is an example of how a stent fabricated from strain inducedaustenite works overlaid with the data from FIG. 11. The displacementaxis represents a small about of bending taking place in the struts ofindividual stent cells. These tiny stent cell displacements could alsorepresent changes in stent diameter. Roman Numerals I that IV show thestent deployment relative to the material properties. Therefore, I showsa prepared stent loaded into the delivery system, II shows the near zerodeployment and expansion force, III shows the fully expanded stent atthe location of complete material transformation, and IV shows how thetransformed stent material can now provide resistance to chronic radialforce.

Surgical Staples

Surgical staples are typically made from 316L and titanium wire havingbeen formed and loaded into delivery magazines. In many applications, adelivery magazine can hold 100 or more staples and they can besimultaneously fired at once. It is desirable to lower the combinedfiring force to push multiple staples from the magazine holder andsimultaneously crimp the wire staples into the traditional B-shapeprofile.

The present application can improve upon multiple staple tiring systemsby reducing the total work and maximum force required to deploy a givennumber of staples and compress those staples into the required B-shapedprofile. The initial low stiffness of the present disclosure allowsengineers to redesign either: (a) surgical staple devices that areeasier for the physician to grasp and tire, or (b) allow engineers todesign surgical staple guns that fire more staples for an equivalentgrasping force.

Medical Expansion Bolt/Bulkhead Connector Applications

it is also possible to utilize the strain induced austenite disclosureto design and fabricate a device that can fit through a small blindhole. Once inserted and mechanically expanded, the device can take onthe new structural shape of a particular design intent that can preventits removal. One particular application is the Vascular Hole Closuredevice in which a vessel wall is sealed from the outside wall. This canbe achieved by inserting a device comprising the material into thevessel hole (resulting from a previous medical procedure) and causing acontrolled deformation deployment means expanding the device in such away that it cannot be removed, and thereby seals the hole. A secondsimilar application in which the material may be useful is the atrialseptal defect device.

It will now be apparent to those skilled in the art that otherembodiments, improvements, details and uses can be made consistent withthe letter and spirit of the foregoing disclosure and within the scopeof this patent, which is limited only by the following claims, construedin accordance with the patent law, including the doctrine ofequivalents.

What is claimed is:
 1. A method for the thermal mechanical treatment ofa shape memory ternary alloy, said alloy comprising a third insolublealloying element, said method performed at temperatures <M_(d) whereinthe hysteresis is widened such that after the treatment is completedA_(s)<A_(s)′ comprising: a) annealing or partially annealing a shapememory alloy through a high temperature solution treatment followed bycooling; b) mechanically straining the alloy under a load whilesimultaneously cooling the material to a temperature around M_(s), toproduce locked martensite thus widening the hysteresis such thatA_(s)<A_(s)′ and retaining a sufficient amount of strain in the strainedelement whereby if controlled deformation is applied the alloy istransformed to the austenitic phase, shifting A_(s)′ to A_(s) byunlocking the martensite; and c) applying controlled deformationsufficient to initiate the unlocking of martensite whereby the alloy istransformed to austenite wherein A_(s)′ is shifted to A_(s).
 2. A methodfor manufacturing a medical device for use in a body lumen, comprising:forming a structural element from the alloy manufactured by the methodof claim 1 into a desired medical device geometry which includes ahollow structure comprising a mesh like pattern having struts and nodes,the structural element being capable of assuming a first position, asecond position and a third position wherein the first position, thealloy has been processed in an austenitic phase and wherein a secondposition the alloy of the first position is cooled and mechanicallycollapsed for deployment into the body lumen, the nodes achieve highlevels of straining which can be retained in the collapsed secondposition as locked martensite and, wherein the third position isachieved after controlled deformation is applied to mechanically expandthe hollow structure thereby releasing the retained strain to transformthe locked martensite nodes into austenite, enhancing the structure'sability to resist radial loads.
 3. The method according to claim 2,wherein the alloy is Ni—Ti.
 4. The method according to claim 2 whereinwhile mechanically collapsing the structure such that the nodes arestrained between about 4% and 25% under a sufficient load whilesimultaneously cooling the material to a temperature to approximatelyM_(s), thus shifting the A_(s) of the highly strained nodes up highersuch that A_(s)<A_(s)′ and retaining about 2% to about 16% strain in thenodes whereby if controlled deformation is applied the nodes willtransform to the austenite, shifting the nodes transformation propertyfrom A_(s)′ to A_(s).
 5. The method according to claim 4, whereby thestructural element incorporates collapsing of the structural element fordeployment and the pre-straining of nodes into a single process step. 6.The method according to claim 5, whereby the structural element ismechanically expanded and nodes are mechanically transformed by theexpansion process from A_(s)′ to A_(s).
 7. The method of claim 2 wherebythe alloy of the second position is stable for insertion into the bodylumen without a restraining cover and A_(s)′>body temperature.
 8. Amethod for manufacturing a medical device for use in a body lumen,comprising: forming a structural element from the alloy manufactured bythe method of claim 1 into a desired medical stent device geometry whichincludes a hollow structure, the structural stent element having aplurality of stent cells and said element being capable of assuming afirst position and a second position wherein in the first position, thealloy has been processed in an austenitic phase and wherein in thesecond position, the alloy of the first position is cooled andmechanically collapsed for deployment into the body lumen, and compriseshigh levels of strain in the stent cells which is retained in thecollapsed second position as locked martensite, wherein after controlleddeformation is applied to mechanically expand the hollow stent structurethereby releasing the retained strain, transforming the lockedmartensite into austenite, enhancing the structure's ability to resistradial loads.
 9. The method according to claim 8, wherein the alloy isNi—Ti.
 10. The method according to claim 8 wherein while mechanicallycollapsing the structure such that the stent cells when collapsed arestrained between about 4% and 25% under a sufficient load whilesimultaneously cooling the material to a temperature to approximatelyM_(s), thus shifting the A_(s) of the highly strained stent cells uphigher such that A_(s)<A_(s)′ and retaining about 2% to about 16% strainin the stent cells whereby if controlled deformation is applied, thestent cells will transform to the austenite, shifting the stent cellstransformation property from A_(s)′ to A_(s).
 11. The method of claim 8whereby the alloy of the second position is stable for insertion intothe body lumen without a restraining cover and A_(s)′>body temperature.